This invention relates to scintillator structures and methods for fabricating such structures. More particularly, this invention relates to a method for distributing the scintillator phosphor in such a way as to enhance the escape of the visible wavelength radiation that would otherwise be dissipated within a scintillator body. Two embodiments of the present invention are disclosed: one in which the phosphor is distributed in a layered structure and another in which the phosphor is dispersed throughout a transparent matrix.
In general, a scintillator is a material which emits electromagnetic radiation in the visible or near-visible spectrum when excited by high energy electromagnetic photons such as those in the x-ray or gamma-ray regions of the spectrum, hereinafter referred to as supra-optical frequencies. Thus, these materials are excellent choices for use as detectors in industrial or medical x-ray or gamma-ray equipment. In most typical applications, the output from scintillator materials is made to impinge upon photoelectrically responsive materials in order to produce an electrical output signal which is directly related to the intensity of the initial x-ray or gamma-ray bombardment.
Scintillator materials comprise a major portion of those devices used to detect the presence and intensity of incident high energy photons. Another commonly used detector is the high pressure noble gas ionization device. This other form of high energy photon detector typically contains a gas, such as xenon, at a high pressure (density), which ionizes to a certain extent when subjected to high energy x-ray or gamma-ray radiation. This ionization causes a certain amount of current flow between the cathode and the anode of these detectors which are kept at relatively high and opposite polarities from one another. The current that flows is sensed by a current sensing circuit whose output is reflective of the intensity of the high energy radiation. Since this other form of detector operates on an ionization principle, after the termination of the irradiating energy there still persists the possibility that a given ionization path remains open. Hence, these detectors are peculiarly sensitive to their own form of "afterglow" which results in the blurring, in the time dimension, of information contained in the irradiating signal as a result of its passing through a body to be examined, as in computerized tomography applications.
As used herein and in the appended claims, the term "light" means those electromagnetic radiations in the visible region of the spectrum and also near-visible wavelengths given off by certain fluorescent materials. Also, as used herein, and in the appended claims, the term "optical" encompasses the same spectral region as the term "light" does.
In general, it is desirable that the amount of light output from these scintillators be as large as possible for a given amount of x-ray or gamma-ray energy. This is particularly true in the medical tomography area where it is desired that the energy intensity of the x-ray be as small as possible to minimize any danger to the patient.
Another important property that scintillator materials should possess is that of short afterglow or persistence. This means that there should be a relatively short period of time between the termination of the high energy radiating excitation and the cessation of light output from the scintillator. If this is not the case, there is resultant blurring, in time, of the information-bearing signal generated, for example, when the scintillator is used to produce tomographic imaging data. Furthermore, if rapid tomographic scanning is desired, the presence of the afterglow tends to severely limit the scan rate, thereby rendering difficult the viewing of moving bodily organs, such as the heart or lungs.
A scintillator body or substance, in order to be effective, must be a good converter of high energy radiation (that is, x-rays and gamma-rays). Typically, present scintillator bodies consist of a phosphor in a powder polycrystalline, or crystalline form. In these forms, the useful light that is produced upon high energy excitation is limited to that which can escape the interior of the scintillator body and that generated in the surface regions. The escape of light is difficult due to the optical absorption resulting from multiple internal reflections, each such reflection further attenuating the amount of light available to external detectors. Thus, it is necessary that not only the phosphors themselves have a good luminescent efficiency but it is also necessary that the light output be available for detection.
In the medical tomography area, where the intensity of x-radiation is modulated by the body through which it passes, and which modulated radiation is then converted into electrical signals, it is important to have x-ray detection devices which have a good overall energy conversion efficiency. For devices with low efficiency, a higher x-ray flux radiation must be applied to produce the same light and electrical output from the overall system. In a medical tomographic context, this means that such a system has a low signal-to-noise ratio.
Typical scintillator phosphors which are used include barium fluorochloride doped with a europium activator (BaFCl:Eu). Other phosphors, for example, include bismuch germanate (Bi.sub.4 Ge.sub.3 O.sub.12), lanthanum oxybromide doped with terbium (LaOBr:Tb), cesium iodide doped with thalium (CsI:Tl), cesium iodide doped with sodium (CsI:Na), calcium tungstate (CaWO.sub.4), cadmium tungstate (CdWO.sub.4), zinc cadmium sulfide doped with silver (ZnCdS:Ag), zinc cadmium sulfide doped with silver and nickel (ZnCdS:Ag,Ni), gadolinium oxysulfide doped with terbium (Gd.sub.2 O.sub.2 S:Tb), and lanthanum oxybromide doped with dysprosium (LaOBr:Dy). Other host-crystal possibilities for phosphors include the selenides of zinc and cadmium, the tellurides of zinc and cadmium, sodium iodide (NaI), and the oxysulfide of lanthanum (La.sub.2 O.sub.2 S).